In human physiology, homeostasis is promoted by a variety of negative feedback loops. In many of these, the concentration of a certain metabolite is measured by a sensing mechanism and hormones are released to restore or maintain the metabolite at optimum concentration. In one such system the metabolite--glucose--is regulated by release of appropriate concentrations of insulin from beta cells in the pancreas; which cells also serve as the glucose sensing mechanism. In diabetes mellitus, insulin production and/or utilization is impaired and exogenous insulin is periodically administered to attempt to restore body homeostasis.
Periodic administration of insulin or of any other drug has the disadvantage that the drug level within the body varies, rising above optimum initially then falling below optimum, resulting in poor maintenance of the patient and inefficient use of the drug. Increasing the number of applications may minimize the adverse effects of high dosages and improve efficiency but also results in higher costs and more inconvenience to the patient.
Therefore, numerous attempts have been made to develop an artificial endocrine pancreas that can respond to changes in plasma glucose by the administration of appropriate quantities of insulin or an insulin antagonist such as glucagon or glucose. Proponents of this concept hoped, in this way, to maintain normoglycemia in diabetic subjects using a negative feedback control system analogous to that utilized by the natural pancreas. Such systems have included some type of a glucose sensor, an electronic control unit, an insulin pump and a drug reservoir.
Pioneering work toward the development of an artificial device for the control of glycemia was reported by Kadish in Am J. Med. Electron., 3, 82 (1964), who used a Technicon Autoanalyzer.RTM. to analyze glucose concentrations in blood drawn through a double lumen catheter at 0.2 ul/hr. By means of electronically controlled servo-mechanisms and a syringe pump, his device was designed to administer insulin if blood glucose exceeded 150 mg/dl or glucagon if it fell below 50 ml/dl. In a trial with a diabetic volunteer, the device responded to hypo- and hyperglycemic challenges by returning blood glucose to the 50-150 mg/dl range, but its sluggish response time (10-11 min) allowed substantial glycemic excursions to occur. Another disadvantage of the device was that it used excessive amounts of blood (288 ml/day).
Pfeiffer et al. in Horm. Metab. Res., 6, 339 (1974) and Albisser et al. in Diabetes, 23, 397 (1974) reported that they were able to essentially normalize plasma glucose concentrations in diabetic volunteers. They used equipment similar to that used by Kadish but with several modifications. An improved version of the Technicon Autoanalyzer.RTM. was used that reduced response time to 4-5 min and reduced blood loss per day of continuous operation to about 70 ml. They also developed control algorithms that altered insulin and insulin antagonist infusion rates according to rates of change in glycemic parameters, as well as static plasma glucose values. These algorithms permitted the machines to anticipate glycemic excursions and respond accordingly. However, all of the devices described above were large extracorporeal units suitable only for acute studies in the hospital setting.
During the early 1970's, Soeldner of the Joslin Clinic and Bessman of the University of South California began work on a miniaturized and simplified form of an artificial pancreas--more appropriately called an artificial beta cell--designed for implantation. As disclosed by Soeldner et al. in Temperal Aspects of Therapeutics, Plenum Press, N.Y. (1973) such a device would consist of a glucose sensor with its accompanying power supply, a computer, a pump, and an insulin reservoir with a self-sealing refill portal. They suggested that the device could include optional features such as telemetered alarm signals to indicate device malfunction or the need to refill the reservoir. Based on the belief that the glucose sensor was the most important component of the system, they began that phase of development first.
In the Technicon Autoanalyzer.RTM.--the equipment used by Albisser and Pfeiffer and colleagues for continuous plasma glucose determination--a continuous stream of blood, drawn from a patient through a double-lumen catheter, is diluted, anticoagulated, and then dialyzed against alkaline potassium ferricyanide. Glucose was then determined colorimetrically. While this method was satisfactory for the extracorporeal units described above, it is too cumbersome for use in an implantable system. Soeldner and colleagues chose, instead, to design an electrochemical sensor based on the property of nobel metals such as platinum to catalyze the oxidation of glucose to gluconic acid. Several electrochemical sensor sub-types can be contructed using this basic principle including fuel cell, polarographic, potentiometric and potentiodynamic systems. In 1973, Chang et al. in Trans. Amer. Soc. Artif. Intern. Organs, 352, 19 (1973) chose the fuel cell type sensor for their initial experiments. A fuel cell is comprised of a nonconsumable catalytic anode and cathode, an electrolyte, and membranes separating the anodic and cathodic environments. The system does not need applied current or a reference electrode, thus reducing the problem of oxide formation and overcoming the problem of reference electrode degradation. Oxide coating on the platinum anode is reduced, but not eliminated by the lack of applied current. The performance of eight of these sensors was tested by subcutaneous implantation in monkeys for up to 117 days. Sensor output, which was transmitted through percutaneous lead wires to an amplifier and a recorder, could not be rigorously correlated with blood glucose values obtained by standard methods. However, the sensor-derived values following meals and during glucose tolerance tests appeared to fall within the expected ranges.
One shortcoming of the electrochemical sensor is its nonspecificity. In addition to glucose, it responds to other monosaccharides, certain amino acids, ethanol, and urea. Since these substances are commonly found in blood and intracellular fluid, their presence can greatly reduce the accuracy of the results obtained by this method.
Like Soeldner, Bessman and his colleagues chose to develop the glucose sensor as the first component of their system. However, the type of unit they chose for study was an enzyme electrode sensor. Like the electrochemical sensor described above, the enzyme electrode catalyzes the oxidation of glucose to gluconic acid and hydrogen peroxide. Unlike the electrochemical sensor, the enzyme electrode is highly specific for glucose. The enzyme electrode glucose sensor, as disclosed by Clark and Lyons in Ann. N.Y. Acad. Sci., 103, 29 (1962) consisted of a glucose oxidase solution sandwiched between semipermeable polymeric membranes. Initially, a pH electrode measured glucose concentration as a function of hydrogen ion concentration, which changed in accordance with the amount of gluconic acid formed. Later sensors potentiometrically measured glucose concentration as a function of oxygen depletion using an oxygen electrode also designed by Clark (Trans. Amer. Soc. Artif. Intern. Organs, 2, 41 (1956)). In a modified enzyme electrode glucose sensor designed by Updike and Hicks, Nature, 214, 986 (1967), gludecose oxidase was bound to a thin layer of polyacrylamide gel. This sensor substantially reduced response times over previous sensor models. It differed from the previous model of Clark by using a polarographic, rather than a potentiometric, oxygen electrode, i.e. by measuring amperage rather than voltage differences.
Bessman and Schultz modified the Clark design further by immobilizing and stabilizing the glucose oxidase by intra- and inter-molecular cross linkages in cloth matrix disks that were cemented over the plastic membrane of a polarographic oxygen electrode (Trans. Amer. Soc. Artif. Intern. Organs, 19, 361 (1973)). This modification extended the useful range of the device up to 400 mg/dl, about twice that of the Updike-Hicks sensor. The pumping system designed by Thomas and Bessman to accompany the glucose sensor consisted of two opposed piezoelectric disk benders, arranged in opposition to form a bellows, connected to a solenoid valve (Trans. Amer. Soc. Artif. Intern. Organs, 21, 516 (1957)). A rectangular wave pulse generator activates the opening and closing of the solenoid valve and, through a step-up transformer, activates the flexing of the disk benders. The system was capable of delivering insulin in pulses of 0.2 .mu.l or less. The delivery rate in this device is a function of the number of pulses per unit time. Prior to 1977, Bessman et al. implanted a pump of this design in an alloxan diabetic dog to deliver insulin into the peritoneal cavity (Excerpta Medica, 413, 496 (1977)). They reported that plasma glucose was maintained within the physiological range for four days using this system.
A needle-type glucose sensor was disclosed by Schichiri et al., in Lancet, 2, 1129 (1982). It is a glucose oxidase sensor similar to those described above. It differs from them by being designed as a small needle that can be inserted in the skin to measure capillary blood glucose. By means of telemetry it can be used to control an implantable insulin pump. However, the device must be replaced at intervals of approximately three days.
Schultz et al. in Diabetes Care, 5, 245 (1982) described an affinity sensor for monitoring various blood metabolites by optical means. Its operating principle is based on competitive binding of the metabolite to be detected and a fluorescent dye-labeled ligand on receptor sites specific for both the metabolite and the labeled ligand. In designing an optical sensor specific for glucose, concanavalin A, a plant glycoprotein that binds glucose, was immobilized on the inside surface of a hollow dialysis fiber. Dextran, a glucose polymer, labeled with fluorescein was sealed within the fiber. The dialysis fiber selected was permeable to glucose but impermeable to dextran. Thus, while the fluorescein-labeled dextran was retained within the chamber, glucose was free to diffuse in and out. A single optical fiber was inserted into the lumen of the hollow dialysis fiber. The optical fiber and associated electronic equipment were used to measure the fluorescence of the free fluorescein-labeled dextran.
As described above, metabolite sensors that have been developed to date are generally designed as components of closed loop feedback control systems that provide infusion of an appropriate drug in response to signals from the sensor. Thus, all of these designs include an electronic interface between the sensor and the drug delivery components.
Although a number of sensor types have performed successfully in laboratory tests, a sensor suitable for longterm implantation has yet to be reported. Electrochemical sensors are relatively nonspecific and tend to respond to substances in blood or body fluid other than the intended metabolite. Enzyme electrode sensors tend to loose their ability to function due to inactivation of the enzyme. All of the sensor types that have been disclosed, including electrochemical, enzyme electrode and affinity sensors, have failed to address the major obstacle to long-term performance of an implantable sensor; namely, the body's invariable attempt to insulate the sensor from the sampling source. Implanted sensors tend to become surrounded with fibrous tissue shortly after implantation subcutaneously or within a body cavity or, if implanted in contact with the blood, tend to become covered with thrombus. Contact between the blood or intracellular fluid and the sensor is thereby impaired. This has constituted the greatest single obstacle to further development of implantable metabolite sensors.
Therefore, a need exists for a system for measuring the concentration of metabolites, such as glucose, but which is effective to retard encapsulation of the metabolite sensor following implantation. A further need exists for an implantable system which can directly alter the amount of drug delivered without an electronic interface between the sensor and the drug delivery components, such as the pump.